Conventionally, radiographic imaging such as X-ray imaging has been widely used for the diagnosis of symptoms in healthcare settings. In particular, as a result of improvements in sensitivity and image quality achieved over its long history, radiographic imaging using an intensifying screen-film system is now used in healthcare settings throughout the world as an imaging system combining high reliability and excellent cost performance. However, this image information is so-called analog image information and does not allow free image processing or instantaneous transmission such as that of digital image information, which continues to develop in recent years.
Radiographic image detectors for detecting digitally processed radiographic images have emerged in recent years, as typified by computed radiography (CR) or flat panel type radiation detectors (flat panel detectors) (FPDs). These devices are able to directly obtain digitized radiographic images and directly display images on an image display device such as a cathode tube or a liquid crystal panel, so image formation on a photographic film is not absolutely necessary. As a result, these X-ray image detectors reduce the need for image formation by silver salt photography and thus dramatically improve the convenience of diagnostic operations at hospitals or medical clinics.
Computed radiography (CR) is presently accepted in healthcare settings as one type of digital technology for X-ray imaging. However, due to insufficient sharpness and spatial resolution, this technology has not reached the image quality level of screen-film systems. Flat panel detectors (FPDs) using thin film transistors (TFTs) are also being developed as even newer types of digital X-ray imaging technology, as described in the paper “Amorphous Semiconductor Usher in Digital X-ray Imaging” by John Rowlands appearing on page 24 of the November 1997 issue of the journal Physics Today or the paper “Development of a High Resolution, Active Matrix, Flat-Panel Imager with Enhanced Fill Factor” by L. E. Antonuk appearing on page 2 of 1997 Vol. 32 of the journal SPIE, for example.
In order to convert radiation to visible light, a scintillator created with an X-ray phosphor having the characteristic of emitting light in response to radiation is used, but it is necessary to use a scintillator with a high luminous efficiency in order to improve the S/N ratio in the imaging of low-dose radiation. The luminous efficiency of a scintillator is typically determined by the thickness of the phosphor layer and the X-ray absorption coefficient of the phosphor. However, as the thickness of the phosphor layer increases, the scattering of emitted light occurs within the phosphor layer, and the sharpness is diminished. By way of contrast, when the sharpness required for image quality is determined, the layer thickness is limited and the light emission luminance is reduced.
For example, in the scintillator panel disclosed in Patent Document 1, a metal reflective layer made of aluminum, an SiO2 film (first dielectric layer), and a TiO2 film (second dielectric layer) are laminated sequentially on a first surface of an aluminum substrate, and the entire laminate is then covered with a protective film for the reflective film over the entire aluminum substrate, to form a support. A scintillator (CsI columnar crystals or the like doped with Ti, Na, or the like) is then provided on the reflective film protective film surface on the TiO2 film of the support. Since the refractive index of the TiO2 film is higher than that of the SiO2 film, the reflectance of light emitted from the scintillator becomes large, and the light emission luminance improves as a result.
Patent Document 2 discloses a radiographic image detector having a scintillator comprising columnar crystals made of thallium-activated cesium iodide, for example, and a photodetector on a support, the radiographic image detector being disposed so that radiation is incident sequentially in the photodetector followed by the scintillator. There is a columnar crystal region on the radiation incident side of the scintillator, and there is a non-columnar crystal region on the side opposite the incident side of the scintillator. In the columnar crystal region, there are columnar crystals yielding good light emission efficiency near the photodetector, and gaps between the columnar crystals serve to guide light and suppress light scattering. As a result, the blurring of the image is suppressed, and light reaching deep areas of the scintillator is also reflected by the non-columnar crystal region, which improves the light emission luminance.
Patent Document 3 discloses a radiographic image conversion panel having an undercoat layer and a phosphor layer on a substrate, wherein the phosphor layer consists of phosphor columnar crystals formed by vapor phase deposition from a phosphor parent compound and an activator, and the orientation of the phosphor columnar crystals based on the X-ray diffraction spectrum of a plane having a plane index of (200), for example, is within the range of from 80 to 100%, regardless of the position of the phosphor layer in the thickness direction. As a result, the derangement of the structure of the phosphor columnar crystals is prevented, and the phosphor suppresses the scattering refraction of light components emitted by X-ray irradiation and propagated in the photoelectric conversion element direction, which increases the light emission luminance of the radiographic image conversion panel.
Even if the reflectance of the reflective layer is improved as in Patent Document 1, the optical transmittance at the base of the columnar crystals is low (Patent Document 2), and the original performance cannot be derived. Therefore, the improvement in the light emission luminance cannot be considered sufficient, and although there is technology which improves the base of the columnar crystals (Patent Document 3), there remains room for improvement.